U.S. patent number 10,143,380 [Application Number 14/412,198] was granted by the patent office on 2018-12-04 for system and method for improved diffuse luminescent imaging or tomography in scattering media.
This patent grant is currently assigned to Lumito AB. The grantee listed for this patent is Lumito AB. Invention is credited to Stefan Andersson Engels, Haichun Liu, Pontus Svenmarker, Can Xu.
United States Patent |
10,143,380 |
Andersson Engels , et
al. |
December 4, 2018 |
System and method for improved diffuse luminescent imaging or
tomography in scattering media
Abstract
A method and system for luminescence molecular imaging or
tomography of a region of interest in a scattering medium is
disclosed. The system comprises a non-linear luminescent marker
material arranged in the scattering medium, one or more light
sources positioned by at least one light source position for
exciting said luminescent marker by excitation light emitted by
said one or more light sources into an excitation volume, a
detector at a luminescent light detection position detecting
luminescence from said luminescent marker due to said excitation
light, wherein said excitation light comprises pulsed excitation
light.
Inventors: |
Andersson Engels; Stefan (Lund,
SE), Xu; Can (Lund, SE), Liu; Haichun
(Lund, SE), Svenmarker; Pontus (Umea, SE) |
Applicant: |
Name |
City |
State |
Country |
Type |
Lumito AB |
Lund |
N/A |
SE |
|
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Assignee: |
Lumito AB (Lund,
SE)
|
Family
ID: |
48741153 |
Appl.
No.: |
14/412,198 |
Filed: |
July 1, 2013 |
PCT
Filed: |
July 01, 2013 |
PCT No.: |
PCT/EP2013/063878 |
371(c)(1),(2),(4) Date: |
December 30, 2014 |
PCT
Pub. No.: |
WO2014/006012 |
PCT
Pub. Date: |
January 09, 2014 |
Prior Publication Data
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Document
Identifier |
Publication Date |
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US 20150196201 A1 |
Jul 16, 2015 |
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Related U.S. Patent Documents
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Application
Number |
Filing Date |
Patent Number |
Issue Date |
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61666899 |
Jul 1, 2012 |
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61771131 |
Mar 1, 2013 |
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Current U.S.
Class: |
1/1 |
Current CPC
Class: |
A61B
5/0071 (20130101); G01N 21/4795 (20130101); G01N
21/6456 (20130101); G01N 21/6428 (20130101); A61B
5/0073 (20130101); G01N 2021/6439 (20130101) |
Current International
Class: |
A61B
5/00 (20060101); G01N 21/47 (20060101); G01N
21/64 (20060101) |
References Cited
[Referenced By]
U.S. Patent Documents
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9012869 |
April 2015 |
Andersson-Engels |
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Foreign Patent Documents
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WO 2010/128090 |
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Nov 2010 |
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WO |
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Other References
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migration, J. Biomed. Opt., vol. 13 (4), pp. 041304-1-041304-10,
Jul. 9, 2008. cited by applicant .
Culver et al., Optimization of optode arrangements for diffuse
optical tomography: A singular-value analysis, Opt. Lett., vol. 26,
No. 10, pp. 701-703, May 15, 2001. cited by applicant .
Gainer et al., Control of Green and Red Upconversion in
Nayf4:Yb3+,Er3+ Nanoparticles by Excitation Modulation, Journal of
Materials Chemistry, vol. 21, No. 46, pp. 18530-18533, XP55122657,
ISSN: 0959-9428, DOI: 10.1039/C1JM13684D, Jan. 1, 2011. cited by
applicant .
Gainer, et al.; Toward the Use of Two-Color Emission Control in
Upconverting NaYF4:Er3+,Yb3+ Nanoparticles for Biomedical Imaging,
Nanoscale Imaging, Sensing, and Actuation for Biomedical
Applications VIII, Proceedings of SPIE, vol. 8231, pp.
823101-1-823101-8, Feb. 1, 2012. cited by applicant .
Gao, In Vivo Cancer Targeting and Imaging with Semiconductor
Quantum Dots, Nature Biotechnology, vol. 22, No. 8, pp. 969-976,
2004. cited by applicant .
International Preliminary Report on Patentability dated Sep. 19,
2014 for PCT Application No. PCT/EP2013/063878, filed on Jul. 1,
2013. cited by applicant .
International Search Report and Written Opinion dated Sep. 5, 2013
for PCT Application No. PCT/EP2013/063878, filed on Jul. 1, 2013.
cited by applicant .
Liu, et al.; Multibeam Fluorescence Diffuse Optical Tomography
Using Upconverting Nanoparticles, Optics Letters, vol. 35, No. 5,
pp. 718-720, Mar. 1, 2010. cited by applicant .
Maestro, et al.; Nanoparticles for Highly Efficient Multiphoton
Fluorescence Bioimaging, Optics Express, vol. 18, No. 23, pp.
23544-23553, Nov. 8, 2010. cited by applicant .
Xu, et al., Fluorescence diffuse optical tomography using
upconverting nanoparticles, Applied Physics Letters, vol. 94(3),
251107, Jun. 23, 2009. cited by applicant .
Xu, et al.; High-Resolution Fluorescence Diffuse Optical Tomography
Developed with Nonlinear Upconverting Nanoparticles, ACS Nano, vol.
6, No. 6, pp. 4788-4795, May 8, 2012. cited by applicant .
Yi, et al., Synthesis, Characterization, and Biological Application
of Size-Controlled Nanocrystalline NaYF.sub.4:Yb,Er
Infrared-To-Visible Up-Conversion Phosphors, Nano Letters, vol. 4,
No. 11, pp. 2191-2196, Oct. 16, 2004. cited by applicant.
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Primary Examiner: Porta; David
Assistant Examiner: Boosalis; Faye
Attorney, Agent or Firm: Knobbe, Martens, Olson & Bear,
LLP
Claims
The invention claimed is:
1. A method of imaging a region in a scattering medium by diffuse
luminescence molecular imaging, the method comprising: providing at
least one non-linear luminescent marker in said scattering medium
at a marker position in said region; exciting said non-linear
luminescent marker by excitation light emitted by one or more light
sources into an excitation volume from at least one light source
position; detecting luminescence from said luminescent marker due
to said excitation light by a detector at a luminescent light
detection position, and wherein said excitation light comprises
pulsed excitation light; and matching pulse characteristics of said
at least one pulse with energy level transitions conditions of said
non-linear luminescent marker to substantially provide for a
desired population of energy levels of said non-linear luminescent
marker related to emission of upconverted light so that said
upconverted light is produced in an efficient manner.
2. The method of claim 1, wherein said pulsed excitation light
comprises at least one pulse of light, and said method further
comprises: exciting said non-linear luminescent marker with a first
pulse; and detecting luminescence from said luminescent marker due
to said excitation light from said first pulse for providing single
pulse luminescence molecular imaging from said first pulse.
3. The method of claim 2, further comprising determining a pulse
length of said pulsed excitation light to be in the range of about
20-200 ms for said single pulse luminescence molecular imaging.
4. The method of claim 1, further comprising determining a pulse
length of said pulsed excitation light to be in the range of about
1-100 ms.
5. The method of claim 1, further comprising varying the power
density of said pulsed excitation light as a function of time,
determining a quantum yield dependence of the luminescence on said
power density, and determining a relative depth coordinate of said
marker position in said scattering medium based on said quantum
yield dependence.
6. The method of claim 5, further comprising determining said
relative depth coordinate based on a derivative of said quantum
yield dependence.
7. The method of claim 5, further comprising exciting in sequence
said non-linear luminescent marker with a first and second pulse
having first and second power densities respectively, and
determining said relative depth coordinate based on a variation in
said quantum yield from said first and second pulses.
8. The method of claim 1, further comprising determining a
dependence of said detected luminescence on the power of said
excitation light for setting a predetermined characteristic of said
pulsed excitation light.
9. The method of claim 1, further comprising: providing movement
between said light source position and said marker position, and
imaging said luminescent marker based on a non-linear dependence of
said detected luminescence on the excitation light intensity and
said light source position in relation to said marker position,
wherein said non-linear dependence is given by the relationship:
L=k*E^x, wherein: E is excitation light intensity in said
excitation volume L is luminescence light intensity from said
luminescent marker k is a positive constant, and x is a positive
number larger than one.
10. The method of claim 1, further comprising: scanning said
excitation light between a plurality of said light source positions
such that said light source position is moved in relation to said
marker position.
11. The method of claim 10, further comprising: detecting said
luminescence for each of said plurality of light source positions,
said luminescence having a total luminescence intensity of said
luminescent marker for each of said plurality of light source
positions, and imaging said luminescent marker by making an image
of said total luminescence intensity for each of said plurality of
light source positions.
12. The method of claim 10, wherein said plurality of light source
positions forms a grid pattern, said luminescence marker having a
projected area on said grid pattern, wherein said projected area is
less than the area covered by said grid pattern, and wherein said
excitation volume is substantially localized to each of said
plurality of light source positions such that said luminescent
marker is partially excited if said light source position overlaps
partially with said projected area.
13. The method of claim 1, further comprising: exciting said
luminescent marker by a first light source having a first
wavelength from a first light source position, and exciting said
luminescent marker by a second light source having a second
wavelength from a second light source position.
14. The method of claim 13, wherein said luminescent marker is
excited by said first and second light sources simultaneously.
15. The method of claim 1, wherein said diffuse luminescent imaging
comprises diffuse luminescent tomography, and/or power-scanning
tomography using a single excitation point.
16. The method of claim 1, further comprising providing metallic
nanostructures at said medium for exposure to said pulsed
excitation light.
17. A method of imaging a region in a scattering medium by diffuse
luminescence molecular imaging, the method comprising: providing at
least one non-linear luminescent marker in said scattering medium
at a marker position in said region; exciting said non-linear
luminescent marker by excitation light emitted by one or more light
sources into an excitation volume from at least one light source
position; and detecting luminescence from said luminescent marker
due to said excitation light by a detector at a luminescent light
detection position, and wherein said excitation light comprises
pulsed excitation light; and time-delaying the detection of the
luminescence.
18. The method of claim 17, further comprising detecting said
luminescence during a time interval succeeding a pulse of said
excitation light.
19. A system for diffuse luminescence molecular imaging of a region
of interest in a scattering medium, said system comprising a
luminescent marker for use in said luminescent molecular imaging of
said scattering medium, wherein said luminescent marker is a
non-linear luminescent marker arranged in said scattering medium,
said system comprising: one or more light sources positioned by at
least one light source position for exciting said luminescent
marker by excitation light emitted by said one or more light
sources into an excitation volume; a detector at a luminescent
light detection position detecting luminescence from said
luminescent marker due to said excitation light, wherein said
excitation light comprises pulsed excitation light; and a
processing unit configured to match pulse characteristics of said
at least one pulse with energy level transitions conditions of said
non-linear luminescent marker to substantially provide for a
desired population of energy levels of said non-linear luminescent
marker related to emission of upconverted light so that said
upconverted light is produced in an efficient manner.
20. The system of claim 19, further comprising a detector unit that
is operable to detect said luminescence during a time interval
succeeding a pulse of said excitation light.
21. The system of claim 19, further comprising a processing unit
operable to determine a pulse length of said excitation light based
on calculation of energy level transitions conditions of said
non-linear luminescent marker.
22. The system of claim 19, further comprising a control unit
operable to vary the power density of said pulsed excitation light
as a function of time, and a second processing unit operable to
determine a quantum yield dependence of the luminescence on said
power density, and to determine a relative depth coordinate of said
marker position in said scattering medium based on said quantum
yield dependence.
23. The system of claim 19, wherein said luminescent marker is a
luminescent biological marker, and said scattering medium is tissue
of a human or animal, said luminescent biological marker being
arranged in said tissue.
24. The system of claim 19, wherein said luminescent marker
comprises nanosized upconverting particles of sodium yttrium
tetrafluoride (NaYF.sub.4), co-doped with either
Yb.sup.3+/Er.sup.3+ or Yb.sup.3+/Tm.sup.3+.
25. The system of claim 19, wherein said luminescent marker
comprises nanosized upconverting particles comprising particles
that are water soluble, and/or particles coated with a structure
that is polar, and/or particles having hydroxyl groups attached the
surfaces of the upconverting particles.
26. The system of claim 19, wherein said luminescent marker has a
protective coating, and/or is biofunctionalized.
27. The system of claim 19, wherein said system is devised for
luminescence molecular tomography.
28. The system of claim 19, wherein said non-linear markers are
attached to an imaging contrast agent for an imaging modality
different from a modality for said luminescent imaging.
29. The system of claim 19, wherein said non-linear marker is
attached to an organic gadolinium complex or gadolinium compound,
which has paramagnetic properties, and wherein said system further
comprises a magnetic resonance imaging (MRI) apparatus for
simultaneous imaging of said region of interest by MRI and
luminescence molecular tomography.
30. The system of claim 19, wherein said excitation light is
provided by a first light source having a first wavelength from a
first light source position, and a second light source having a
second wavelength from a second light source position.
31. The system of claim 30, wherein said excitation light is
provided by said first and second light sources simultaneously.
32. A system for diffuse luminescence molecular imaging of a region
of interest in a scattering medium, said system comprising a
luminescent marker for use in said luminescent molecular imaging of
said scattering medium, wherein said luminescent marker is a
non-linear luminescent marker arranged in said scattering medium,
said system comprising: one or more light sources positioned by at
least one light source position for exciting said luminescent
marker by excitation light emitted by said one or more light
sources into an excitation volume; a detector at a luminescent
light detection position detecting luminescence from said
luminescent marker due to said excitation light, wherein said
excitation light comprises pulsed excitation light; and a
processing unit configured to time-delay the detection of the
luminescence.
Description
FIELD OF THE INVENTION
This invention pertains in general to the field of
photoluminescence imaging or photoluminescence tomography of
absorbing and scattering media, and in particular to a method and
system for such imaging.
BACKGROUND OF THE INVENTION
An example of a scattering medium which is of interest for
photoluminescence imaging (in short luminescence imaging) or
photoluminescence tomography (in short luminescence tomography) is
biological tissue. Tissue optics is a field devoted to study the
interaction of light with such tissue. Over the last decades, the
field has grown rapidly. With increasing knowledge of the
light-tissue interaction, the interest in applying tissue optics as
a diagnostic tool is also emerging, reaping the fruits from the
fundamental research.
An area in tissue optics, which the present disclosure is partly
dealing with, is photoluminescence imaging including
photoluminescence tomography, which are non-invasive approaches for
in-vivo imaging of humans or animals. These imaging approaches are
luminescence-based and require an external source of light for
excitation of luminescent biological markers.
Photoluminescence is a process in which a substance absorbs photons
and then re-radiates photons. A specific form of luminescence is
fluorescence, where typically emitted photons are of lower energy
than those used for illumination. Thus, in fluorescence, the
fluorescent wavelength is Stokes shifted to a longer wavelength
with reference to the wavelength of the illuminating light.
Fluorescent imaging is known and can, for example, be used to study
biological responses from drugs in small animals over a period of
time, without the need to sacrifice them.
Shimomura, Chalfie and Tsien were rewarded with the Nobel prize in
2008 for discovering and developing the green fluorescent protein,
which has become a very important fluorescent marker.
However, hitherto, fluorescence molecular imaging and tomography
systems for diffuse luminescent imaging or tomography in absorbing
and scattering media suffer from a number of drawbacks. They have
for instance a low resolution or contrast, which makes diagnostic
tasks based on the imaging results difficult.
Further problems with previous techniques are low quantum yield,
shallow imaging depths, long data acquisition times, and thermal
side effects.
Thus, there is a need for an improved diffuse luminescent imaging
or luminescent tomography system and method which in particular
allow for increased effectiveness by improving the aforementioned
drawbacks.
SUMMARY OF THE INVENTION
Accordingly, embodiments of the present invention preferably seek
to mitigate, alleviate or eliminate one or more deficiencies,
disadvantages or issues in the art, such as the above-identified,
singly or in any combination by providing a system, a method, and
uses according to the appended patent claims.
According to a first aspect of the invention, a method of imaging a
region in a scattering medium by diffuse luminescence molecular
imaging is provided. The region comprises at least one luminescent
marker arranged in the scattering medium at a marker position,
where the luminescent marker is a non-linear luminescent marker.
The method comprises exciting the luminescent marker by excitation
light emitted by one or more light sources into an excitation
volume from at least one light source position, detecting
luminescence from the luminescent marker due to the excitation
light by a detector at a luminescent light detection position,
wherein the excitation light comprises pulsed excitation light.
According to a second aspect of the invention, a system for diffuse
luminescence molecular imaging of a region of interest in a
scattering medium is provided. The system comprises a luminescent
marker for use in the luminescent molecular imaging of the
scattering medium, where the luminescent marker is a non-linear
luminescent marker arranged in the scattering medium. The system
comprises one or more light sources positioned by at least one
light source position for exciting the luminescent marker by
excitation light emitted by the one or more light sources into an
excitation volume. The system comprises a detector at a luminescent
light detection position detecting luminescence from the
luminescent marker due to the excitation light, wherein the
excitation light comprises pulsed excitation light.
In embodiments the luminescent marker is comprised in a group of
non-linear luminescent markers configured to upconvert incoming
light of an illumination wavelength, such that luminescence occurs
at a luminescence wavelength that is shorter than said illumination
wavelength when said luminescent marker is illuminated with said
incoming light.
The luminescent marker is in certain embodiments a biological
luminescent marker.
According to another aspect of the invention, a use of a system of
the second aspect of the invention is provided for luminescence
imaging or tomography of tablets and/or for diffuse optical
imaging, and/or photodynamic therapy, and/or remote activation of
biomolecules in deep tissues, and/or single-shot deep tissue
imaging, and/or for in-vivo or in-vitro luminescence imaging or
luminescent tomography of a small animal, and/or for functional
diagnostics, such as cancer diagnostics, by said luminescence
imaging or luminescent tomography, and/or superresolution
microscopy comprising stimulated emission depletion (STED) or
single-molecule detection using said non-linear luminescent marker
as probe.
In an embodiment, the non-linear markers are attached to an imaging
contrast agent for another imaging modality. For instance a
non-linear marker is attached to a contrast agent for imaging with
a conventional imaging modality, such as Magnetic Resonance Imaging
(MRI), X-Ray, etc. In a specific embodiment, a non-linear marker is
attached to an organic gadolinium complex or gadolinium compound,
which has paramagnetic properties.
Further embodiments of the invention are defined in the dependent
claims, wherein features for the second and subsequent aspects of
the invention are as for the first aspect mutatis mutandis.
Some embodiments provide for increased emission intensity.
Some embodiments provide for increased resolution in diffuse
luminescence molecular imaging and in fluorescence molecular
tomography.
Some embodiments provide for determination of distribution of
ingredients in tablets. For instance, a non-linear luminescent
marker or fluorophore may be attached to an active ingredient in a
tablet. The spatial distribution of the active ingredient may thus
advantageously be determined.
Some embodiments provide for enhanced contrast in medical magnetic
resonance imaging, when non-linear markers are used as an MRI
contrast agent. At the same time, luminescence imaging or
tomography may be made, providing for functional diagnostic
information combined with high resolution MRI of one and the same
region of interest and in-vivo.
Some embodiments provide for increased quantum yield when using
upconverting nanoparticles.
Some embodiments provide for single-shot deep tissue imaging.
Some embodiments provide for large imaging depths and short data
acquisition times.
Some embodiments provide for suppressing of thermal side effects of
the excitation light.
Some embodiments provide for diffuse optical imaging, photodynamic
therapy and remote activation of biomolecules in deep tissues.
Some embodiments provide for a background-free signal.
It should be emphasized that the term "comprises/comprising" when
used in this specification is taken to specify the presence of
stated features, integers, steps or components but does not
preclude the presence or addition of one or more other features,
integers, steps, components or groups thereof.
BRIEF DESCRIPTION OF THE DRAWINGS
These and other aspects, features and advantages of which
embodiments of the invention are capable of will be apparent and
elucidated from the following description of embodiments of the
present invention, reference being made to the accompanying
drawings, in which
FIG. 1 is a Jablonski diagram;
FIGS. 2 a)-c) are schematic illustrations of a) radiative and
nonradiative energy transfer; b) Resonant and nonresonant energy
transfer; and c) Comparison of ETU (left) and ESA (right)
upconversion;
FIG. 3A is a schematic illustration of an upconversion processes in
the Yb.sup.3+--Tm.sup.3+ ion pair of a upconversion nanocrystal;
FIG. 3B is a graph showing the emission spectrum for the
upconversion nanocrystals of FIG. 3A and the excitation power
density dependence of the upconversion emission;
FIGS. 4a-d are schematic illustrations of planar imaging
implementations, namely (a)-(b) setup used for fluorophore imaging
(epi-fluorescence); (d) a setup to be used for fluorophore
reconstruction in transillumination; and (c) another setup for
fluorescence diffuse optical tomography.
FIGS. 5a-c are schematic illustrations of the difference between
fluorescence imaging with linear and non-linear fluorophores;
FIG. 6 is a graph showing the normalized singular-value
distribution of a weight matrix W, for single-beam excitation and
combined single-beam excitation and dual-beam excitation.
FIGS. 7A-B are three-dimensional reconstructions of upconverting
nanoparticles, using (10A) only single-beam images, and using (10B)
both single-beam and dual-beam images.
FIG. 8 shows upconversion spectrum of NaYF4:Yb3+, Tm3+
nanoparticles under excitation of 975 nm;
FIG. 9 shows power dependencies of near infrared, red and blue
upconversion emission bands of NaYF4:Yb3+, Tm3+ nanoparticles under
excitation of 975 nm determined according to an embodiment of the
invention;
FIG. 10 shows quantum yields of near infrared, blue, red
upconversion emission bands at various power densities determined
according to an embodiment of the invention;
FIG. 11 shows upconversion spectra of NaYF4:Yb3+,Tm3+ nanoparticles
under excitation of CW and pulse excitation (with identical average
power) according to an embodiment of the invention;
FIG. 12 shows upconversion emission spectra (normalized at 800 nm)
under CW excitation and pulse excitation with different pulse width
(1 ms, 5 ms, 10 ms) according to an embodiment of the
invention;
FIG. 13 shows upconversion spectra under excitation of CW and
pulses (with the same average power of 31.04 mW) according to an
embodiment of the invention;
FIG. 14 shows the gain of signal at 800 nm under pulse excitation
with 5 and 10 ms pulse widths under different average excitation
power according to an embodiment of the invention;
FIGS. 15a-f and FIG. 16 show the influence of excitation power on
the gain at 800 nm;
FIGS. 17a-c and 18a-b show the influence of excitation power on the
gain at 800 nm;
FIG. 19a-c show luminescence images taken at 800 nm under the
excitation of; a CW laser diode with a power of 100 mW (a); a pulse
laser (square wave, pulse width 5 ms, period 250 ms) with an
average power of 100 mW (b); and a pulse laser (square wave, pulse
width 10 ms, period 500 ms) with an average power of 100 mW
(c);
FIG. 20 illustrates excitation light comprising pulsed excitation
light according to embodiments of the invention;
FIG. 21 illustrates luminescence and delayed detection of
luminescence following pulsed excitation light according to
embodiments of the invention;
FIGS. 22-25 illustrate relative depth coordinates of markers and
the determination thereof according to embodiments of the
invention;
FIGS. 26a-b illustrates signal gain versus pulse width and average
power density respectively;
FIGS. 27a-d illustrates luminescence from a marker following pulsed
excitation according to embodiments of the invention (b, d) and
luminescence from continuous wave (CW) excitation (a, c);
FIG. 28 illustrates a schematic flow-chart of a method according to
embodiments of the invention;
FIGS. 29a-c illustrates simulated quantum yield (QY) versus time
and average power density; and
FIG. 30 illustrates the upconversion signal gain versus power
density.
DESCRIPTION OF EMBODIMENTS
Some embodiments of this disclosure pertain to an area within the
aforementioned tissue optics dealing with diffuse luminescence
imaging and tomography. For most visible wavelengths, light does
not penetrate more than a few millimeters into tissue. But in the
diagnostic window (wavelength 600 to 1600 nm), the light
penetration is sufficient to allow imaging through up to several
centimeters. This opens up the possibility of imaging fluorescent
contrast agents deep in tissue. Previous techniques limit the depth
of imaging due to low quantum yield, which also lead to long
acquisition times, noise and thermal side effects.
Experiments on tissue phantoms, with realistic optical properties,
were performed, and it was shown that it is possible to improve
these aforementioned factors according to the below disclosure of
the embodiments of the present invention.
It has previously been shown, in WO 2010/128090, which discloses a
system, a method, and non-linear luminescent markers for diffuse
luminescent imaging or tomography that contrast and resolution of
such imaging can be improved.
Several applications within biomedical imaging of the fluorescence
imaging or tomography are described below.
Other applications are provided in non-biological areas. Examples
for such areas are luminescent imaging or tomography for material
testing, including quality control of tablets, filters for liquids
or gases through which flows a medium with non-linear markers,
etc.
In the context of the present application and embodiment of the
invention, fluorescence imaging represents all types of imaging of
luminescence. Also, any imaging or tomography discussed is in
highly scattering media, traditionally providing poor resolution
due to the diffuse character of the light detected. Embodiments of
the present invention advantageously improve quantum yield,
contrast and resolution of such luminescent imaging, including in
luminescent tomography.
Specific embodiments of the invention will now be described with
reference to the accompanying drawings. This invention may,
however, be embodied in many different forms and should not be
construed as limited to the embodiments set forth herein; rather,
these embodiments are provided so that this disclosure will be
thorough and complete, and will fully convey the scope of the
invention to those skilled in the art. The terminology used in the
detailed description of the embodiments illustrated in the
accompanying drawings is not intended to be limiting of the
invention. In the drawings, like numbers refer to like
elements.
Below, an overview of the fundamentals of fluorescence imaging,
tissue optics and non-linear markers, such as upconverting
nanocrystals are given, followed by a description of results from
experiments and simulations. More details are given in WO
2010/128090.
Fluorescence Contrast
The process of light emission from a fluorescing molecule
(fluorophore) can be described in a Jablonski diagram, see FIG. 1.
FIG. 1 shows a Jablonski diagram showing the various decay paths
from an excited state of a molecule. In the lower part of the
figure, a fluorescence spectrum from haematoporphyrin in ethanol is
shown. The abbreviations are: Sn: singlet states; Tn: triplet
states; Abs: absorption; Sc: scattering; IC: internal conversion;
F: fluorescence; IX: intersystem crossing; P: phosphorescence; A:
transfer to other molecules. Also the approximate time-scale for
some processes is shown down right in FIG. 1, as lifetimes (LT),
also denoted .tau..
If an incoming photon has an energy that corresponds to the gap
between two energy bands in the molecule, it can be absorbed. The
photon energy will thereby be used for excitation of the molecule
to the higher energy band. Excited states are unstable and the
molecule will return to the ground state. The deexcitation may
follow a number of different pathways, as illustrated in FIG. 1.
The labeled levels are electronic levels, corresponding to the
energy levels of atoms. S0, S1, etc. are singlet states for which
the sum of the electron spin quantum numbers is zero, while T0, T1,
etc. are triplet states for which the spin of one electron has
changed sign. For large molecules the intervals between the levels
are very small and the states overlap due to molecular
interactions. When a photon is absorbed by a molecule it will not
necessarily excite the molecule to the lowest vibrational level in
the excited electronic level, but more likely to a higher
vibrational state. This is a result of the Franck-Condon principle
stating that during the rapid (10.sup.-15 s) absorption process,
the atoms do not change their location in the vibrational motion.
When a molecule is excited to a high energy level, a rapid
relaxation to the lowest rotational-vibrational state of S1 will
follow. The short time scale (10.sup.-12 s) of this relaxation is
due to the high density of rotational vibrational levels. From S1
the molecules can proceed to the state S0 through radiationless
kinetic interactions. This is called internal conversion (IC).
Alternatively, the de-excitation may result in the emission of a
photon and this process is called fluorescence. Since the
transition may be terminated in any of the rotational-vibrational
states of S0, the energy of the different photons will not have a
distinct value, but rather a broad distribution. Thus, a
fluorescence spectrum from a molecule will be broad, most often
without any significant structures. The form of the spectrum will
reflect the probability of transitions to the lower levels (S0). In
the lower part of FIG. 1 the fluorescence spectrum of
haematoporphyrin, which is a tumour marker, or photosensitizer, and
will be discussed later on, is shown. Once the pathway
absorption-IC-fluorescence is completed, the molecule is back in
its original state and configuration. Hence, the fluorescence
process is non-destructive and reversible, which is an advantage
in, for instance, medical diagnostics.
Several other paths are possible for the excited molecule, such as
energy transfer to other molecules, electron transfer, excimer
formation and excitation to repulsive states leading to molecular
dissociation. These processes are indicated with an A in FIG.
1.
Many fluorescent molecules have one important feature in common,
that is the unbroken chain of conjugated double bonds, i.e. every
second bond is a double bond. The structure of haematoporphyrin is
an example for this (not shown). This is a fluorescent molecule
used for fluorescence diagnostics and photodynamic therapy of
tumours.
Fluorescence Imaging
In contrast to point monitoring devices, Fluorescence imaging
systems can detect a fluorescence signal in large number of points.
Thus, a two-dimensional image of an area of interest is created. A
typical system comprises a camera together with a tunable filter,
see FIG. 4a. A similar setup in transillumination is schematically
illustrated in FIG. 4c. With a tunable filter the wanted detection
wavelengths can easily be selected and a spectral resolution of
approximately 20 nm wide may be achieved.
Fluorescence Imaging with Non-Linear Fluorophores
A particularly interesting subsection of fluorescence imaging is
that of using non-linear fluorophores. In the context of the
present application, a "non-linear marker" is a luminescent marker,
wherein a luminescence (L) of the marker is not linearly dependent
on the luminous flow of excitation light (E). Non-linear markers
thus have a luminescence according to: L=k*E^x, wherein x>1, and
wherein k is a positive constant. The non-linear markers may also
have a luminescence according to the following relationships:
L=k*E^x+b, L=k(E)*E^x+b, L=k(E)*E^x+b(E), or L=k*E^x+b(E), where k
and b are material constants that are either constant or depending
on the local field of excitation light (E), i.e. for k(E) and b(E).
In comparison to conventional luminescence imaging, non-linear
markers (or fluorophores) may thus require more than one photon for
excitation. This drastically decreases the excitation volume and
provides a more localized excitation point. In this manner,
contrast and resolution of luminescent imaging is improved, as is
demonstrated below. In more detail, contrast and resolution of
diffuse light in luminescent imaging of absorbing and scattering
media is improved. Embodiments of the present invention take
advantage of this effect.
To illustrate the difference between fluorescence imaging with
linear and non-linear fluorophores, reference is made to FIG. 5a-c.
FIG. 5a illustrates a linear fluorescence image in gray-scale. Each
pixel (705) corresponds to one excitation point (704) in a grid
pattern (701). FIG. 5b illustrates an image obtained with a
two-photon, non-linear fluorophore, i.e. non-linear luminescent
marker (702). In FIG. 5c the fluorophore (702) is shown in red
(larger circle) (703), and the black dots (704) indicate the points
of excitation in the grid pattern (701). The circle (703)
corresponds to the projected image of the marker (702) on the grid
pattern (701). The excitation points (704) corresponds to the
positions of the light source, i.e. laser (503), when scanning the
luminescent marker (702). It can clearly be seen that using the
non-linear fluorophore increases contrast and resolution of the
fluorescent image. In particular, when the light source is in the
position marked as 706 in FIG. 5c, close to the marker (702) or
corresponding projected image (703) of the marker (702) on the grid
pattern (701), the excitation volume is sufficiently small and
localized to the light source position (706) for the non-linear
marker, such that no luminescence is detected in the corresponding
pixel (708) in FIG. 5b. For the linear fluorescence image in FIG.
5a, the corresponding pixel (707) receives luminescence due to the
increased excitation volume in the scattering media. The two-photon
non-linear dependence provides the narrow photon-density of the
excitation volume. Thus, imaging the marker (702) based on the
non-linear dependence of the detected luminescence on the
excitation light intensity, the resolution may be increased.
Non-linear fluorophores require in general higher excitation
intensities compared to linear fluorophores and some non-linear
fluorophores even require coherent excitation. In scattering media,
high intensities are difficult to achieve, since light cannot be
focused, but rather spreads in every direction. This makes some
non-linear fluorophores more suitable for fluorescence imaging in
scattering media as compared to others. The fluorophores need to
have an exceptionally high yield, and they may not require coherent
excitation. Up-converting nanoparticles are one such non-linear
fluorophore with high yield and non-coherent excitation.
Due to the quadratic dependence of the emitted fluorescence in e.g.
up-converting nanocrystals, the fluorescence tomography is
improved.
Upconversion
Upconversion is a non-linear process that occurs when two or more
photons are absorbed and a photon of higher energy, than those of
the incoming photons, is released.
The process is for instance observed in materials containing a
meta-stable state that can trap one electron for a long time,
increasing the interaction-probability with another arriving
photon.
In some embodiments, luminescent markers in form of solids doped
with different rare earth ions are used to obtain upconversion.
Upconversion can happen due to numerous processes, which impact the
upconversion process differently depending on the ion pairs and the
excitation intensities.
Some upconversion processes are illustrated in FIGS. 2 a)-c).
Some of the processes involve energy transfer between ions. This
energy diffusion, can be radiative or non-radiative, resonant or
non-resonant.
Furthermore, Energy Transfer Upconversion (ETU) and Excited-State
Absorption (ESA) processes are illustrated in FIG. 2c on the left
respectively on the right of the Figure. Excited state absorptions
happen when an ion, being in an excited state, absorbs one more
photon.
Nanosized Upconverting Crystals
Nanosized upconverting particles are for instance lanthanide doped
oxides (Y.sub.2O.sub.3), which are easy to fabricate.
Other nanosized upconverting particles are for instance fluorides,
which have higher efficiencies than Y.sub.2O.sub.3. The higher
efficiencies can be explained by the low phonon energies in
fluorides, which lower the probability for non-radiative decay.
Further nanosized upconverting particles are for instance made of
sodium yttrium tetrafluoride (NaYF.sub.4), co-doped with either
Yb.sup.3+/Er.sup.3+ or Yb.sup.3+/Tm.sup.3+.
NaYF.sub.4 can crystallize in two phases, cubic or hexagonal,
called .alpha.-NaYF4 and .beta.-NaYF4, respectively. The
upconverted luminescence from the .beta.-phase material is
approximately one order of magnitude higher compared to the
upconverted luminescence from the .alpha.-phase. The non-linear
fluorophores, such as the upconverting nanoparticles may also be
biofunctionalized, giving them, for example, tumor seeking
abilities.
The non-linear fluorophores may be water soluble, allowing for easy
administration in certain applications, such as in solutions for
intravenous, peroral, or enteral administration.
A way to provide upconverting nanoparticles as water soluble, is to
coat the particles with a structure that is polar. Coatings may for
instance be made of polymers or silica. Both synthetic polymers,
for example, Polyethylene glycol (PEG), and natural polymers may be
used for the coating. These polymers are stable in biological
environments and do not interfere with the optical properties of
the nanocrystals in any significant negative way.
Water soluble upconverting nanoparticles may be provided without
coatings. Hydroxyl groups may be attached to the surfaces of the
upconverting nanoparticles, either by chemical bonds or physical
absorption. Hydroxyl groups are by definition formed by covalent
binding, and the final structure has polar properties.
Functionalization
Functionalization of the upconverting nanoparticles may be made in
similar ways as functionalizing quantum dots, such as described in
X. Gao et. al., In vivo cancer targeting and imaging with
semiconductor quantum dots, Nature Biotechnology, 22, 8:969-976,
2004, which is incorporated herein in its entirety for all
purposes. In Gao et. al. methods are described that are applicable
on upconverting rare-earth doped nanoparticles.
The upconverting nanoparticles used in an embodiment in this
disclosure were NaYF.sub.4-crystals prepared according to the
method described in G. Yi et. al., Synthesis, characterization, and
biological application of size-controlled nanocrystalline
NaYF.sub.4:Yb,Er infrared-to-visible up-conversion phosphors. Nano
Letters, 4, 11:2191-2196, 2004, doped with a combination of
Yb.sup.3+ and Tm.sup.3+. The energy diagrams for the two ions are
shown in FIG. 3A. FIG. 3A is a schematic illustration of
upconversion processes in the Yb3+/Tm3+ ion pair. Nonradiative
upconverting processes are illustrated with dashed arrows and
non-radiative decays are omitted for clarity. FIG. 3B is a graph
showing the emission spectrum for these upconverting nanoparticles.
The blue emission line at 477 nm is only visible for higher pump
intensities. The pump-power dependence of the 800 nm line was
measured to be quadratic using low intensities, as seen in the
inset of FIG. 3B, showing intensity (I) on the x-axis and counts
(C) on the y-axis and where the slope (S) of the fitted line (401)
equals 2.
In an embodiment, the non-linear markers are attached to an imaging
contrast agent for another imaging modality. For instance a
non-linear marker is attached to a contrast agent for imaging with
a conventional imaging modality, such as Magnetic Resonance Imaging
(MRI), X-Ray, etc. In a specific embodiment, a non-linear marker is
attached to an organic gadolinium complex or gadolinium compound,
which has paramagnetic properties. When used as an MRI contrast
agent, contrast is enhanced in medical magnetic resonance imaging.
At the same time, luminescence imaging or tomography may be made,
providing for functional diagnostic information combined with high
resolution MRI of one and the same region of interest and
in-vivo.
Other applications are provided in non-biological areas. Examples
for such areas are luminescent imaging or tomography for material
testing, including quality control of tablets, filters for liquids
or gases through which flows a medium with non-linear markers,
etc.
System Setup Examples
Systems for diffuse luminescence molecular imaging are shown
schematically in FIGS. 4a-d. FIGS. 4a-b are schematic illustrations
of setups for fluorophore imaging (epi-fluorescence); and FIG. 4c
is a setup for fluorophore reconstruction in transillumination
which and can be used for simulations of FMT using non-linear
fluorophores and traditional fluorophores. In the latter case the
simulated tissue phantom may be modeled as a semi-infinite cylinder
(510) having uniformly spaced source-detector points (509) around
one plane of the geometry.
A tissue phantom (501) may consist of a solution of intralipid ink
with optical properties determined by a suitable system (500, 600),
such as time-of-flight spectroscopy system, frequency domain
system, or other imaging system in the steady state and time or
frequency domain The fluorophores (502) may be contained in
capillary tubes with inner diameters of 2.4 mm. The concentrations
of the fluorophores may be chosen 1 wt % for the nanoparticles and
1 .mu.M for traditional downconverting fluorophores of the type
DY-781 in comparative studies. The concentration of the
nanoparticles can be chosen to have a reasonable correspondence
with studies using quantum dots, namely a concentration of 1 wt
%.
Using step motors, a fiber coupled laser (503) may be raster
scanned. The positions of the laser in the raster scan may be
described by a grid pattern (701) as shown in FIG. 5. An image may
be acquired for each scanned position with an air cooled CCD (504)
camera sitting behind two dielectric band pass filters centered at
800 nm. FIG. 4c shows a raster scanning setup (507) where the laser
is scanning the tissue phantom (501) from a below position (505).
The CCD (504) may capture one image for every position (506) of the
laser. The positions (506) describes a grid pattern (508) similar
to the grid pattern (701) in FIG. 5. For each position (506) of the
laser, the emitted fluorescence from the entire side of the phantom
(501), i.e. the total luminescence intensity, may be measured and
summed to make up one pixel in the resulting image. Hence the
number of pixels in the image may be given by the number of
excitation positions (506) and not by the number of CCD pixels. The
resolution may thus be determined by the photon-density of the
excitation light from the laser light source (505), and not by the
photon-density of the fluorescence emission light. In this way,
because the two-photon photon-density in the excitation volume is
more narrow than the single-photon photon-density, the resolution
could be increased. When summing the total luminescence intensity a
threshold value may be applied to the detected luminescence. In
this way resolution may be increased. For example, only if the
luminescence intensity is above a defined threshold it will be
added to the total luminescence intensity. The threshold may be
defined as a value in the CCD (504), for example if the
luminescence intensity is below 30% of a peak value it will be
discarded, as it might be considered as a background signal.
Further, if the resulting total luminescence for a pixel, or
position (506) of the laser, is below another threshold value it
may be considered as background signal and removed. Alternatively,
the quadratic intensities of the luminescence signal may be summed.
In this way the resolution may be further increased. For example,
the luminescence intensity detected by the CCD (504), which may
have relative value between 0 and 1 by definition of a peak
intensity value in the CCD, may be multiplied with itself before
added to the total luminescence intensity for the current pixel or
position (506). Further, the total luminescence intensity may be
multiplied with itself for each pixel or position (506). Using the
scanning imaging technique, each pixel in the image may correspond
to the fluorescence induced by a single excitation point, i.e.
light source position (506).
FIG. 4b schematically illustrates a system 600 for diffuse
luminescence molecular imaging according to an embodiment of the
invention. The system 600 comprises a luminescent marker 502 for
use in said luminescent molecular imaging of said scattering
medium, wherein the luminescent marker is a non-linear luminescent
marker arranged in the scattering medium. The system 600 comprises
further one or more light sources 503 positioned by at least one
light source position 505, 506, for exciting the luminescent marker
by excitation light emitted by said one or more light sources into
an excitation volume, and a detector 504 at a luminescent light
detection position detecting luminescence from the luminescent
marker due to said excitation light, wherein said excitation light
comprises pulsed excitation light. Hence, the system 600 is adapted
for diffuse luminescence molecular imaging of a region of interest
in a scattering medium by pulsed excitation light, which provides
for improved quantum yield, less thermal side effects due to less
heating of the medium, deeper imaging depths and shorter
acquisition times. The system 600 will be described further in the
below disclosure in relation to enhancing upconversion emission by
pulse excitation and single pulse imaging with pulsed excitation
light.
Multi-Beam Fluorescence Diffuse Optical Tomography Using
Upconverting Nanoparticles
Additionally, this disclosure demonstrate a method in Fluorescence
diffuse optical tomography to exploit the unique nonlinear power
dependence of upconverting nanoparticles to further increase the
amount of information in a raster-scanning setup by including
excitation with two beams simultaneously. It was found that the
increased information led to more accurate reconstructions.
Fluorescence diffuse optical tomography (FDOT) is a relatively new
modality which seeks to reconstruct the spatial distribution of the
concentration of fluorescent probes inside turbid material. As an
imaging tool, it has a good prospect in biomedical studies to
image, for example, tumors, proteases, and drug effects.
FDOT has numerically very ill-posed issues. In this issue, the
quality of the reconstructions for the fluorescent target is
directly determined by the amount and quality of fluorescence
information obtained from boundary measurements. Instrumental noise
and tissue autofluorescence are the main perturbations of the
measurements, resulting in poor signal quality, and can cause
severe artifacts in the reconstructed results. In order to overcome
this, one could, for example, employ low-noise equipment, use
background subtraction or spectral unmixing. However, such methods
cannot resolve all issues, since they essentially are only
utilizing the present information in a better way rather than
adding new constraints for the reconstructions, i.e., adding new
independent information, which is critical to improve the quality
of the reconstructions.
In a noncontact CCD-based FDOT system, one preferred way to gain
more information is by increasing the number of excitation
positions. However, in order to keep the intensity of the
excitation beam within reasonable levels, there is a limit on the
minimum size of the excitation beam. This implies a practical upper
limit to the highest excitation-position density, since distinct,
i.e., non-overlapping, excitation positions are desired for
reconstructions. It is also possible to employ an anatomical
imaging modality such as magnetic-resonance imaging to provide
a-priori structural information. However, this is at the cost of
significantly increased complexity and reduced flexibility of the
system.
In this disclosure, we present an approach to exploit the quadratic
power dependence of upconverting nanoparticles to gain additional
information by utilizing two beams simultaneously for excitation in
FDOT. The effect of the images taken with dual-beam excitation
(named type-D images) on the reconstructions of the nanoparticle
number density distribution, n, is demonstrated. In addition,
comparisons of reconstructed results between the linear Rhodamine
6G and the quadratic upconverting nanoparticles are made.
The excitation and emission fields can be modeled by two coupled
diffusion equations [Ref. 1]. For quadratic fluorophores, the
fluorescence signal detected at a fixed detector position under
excitation of the k:th beam can be described by the forward model
(1);
.GAMMA..times..times..function..times..function..function..function..time-
s..DELTA..times..times. ##EQU00001## where N denotes the number of
voxels, r.sub.s,d,i denotes the coordinates for source, detector,
and voxel, respectively, and; .DELTA.V.sub.i is the volume of voxel
i. The forward solution of the excitation light is represented by;
[U.sub.e(r.sub.s.sub.k, r.sub.i].sup.2 When exciting the medium
using two beams simultaneously, the detected signal is given by
(2);
&.times..times..times..times..function..times..function..function..functi-
on..function..times..DELTA..times..times..times..GAMMA..GAMMA..times..time-
s..times..function..times..function..times..function..times..function..tim-
es..DELTA..times..times. ##EQU00002## which reveals the involvement
of cross-terms. In a raster-scanning setup (500, 507), if two
images are taken sequentially with one excitation beam scanning
over two positions (named type-S images), and a third image is
taken with two-beam excitation (type-D) above the previous two
positions, the involvement of cross-terms implies that the type-D
image cannot be obtained by any mathematical manipulation from the
existing type-S images, indicating that it is independent and
contains additional information. However, for linear fluorophores,
e.g., Rhodamine 6G, the type-D image is only linear combinations of
the existing type-S images, and will not add more constraints for
the inverse problem. For nonlinear fluorophores, it is deduced that
Eq. (2) can be generalized to include more simultaneous excitation
beams. The significance of the measurements with dual-beam
excitation in the reconstructions was confirmed by the
singular-value analysis of the weight matrix, W, whose elements are
given by (3) [Ref. 1];
W.sub.(s,d),i=U.sub.f*(r.sub.d,r.sub.i)[U.sub.e(r.sub.s,r.sub.i)].sup-
..gamma..DELTA.V.sub.i, (3) with; .gamma.=2 for quadratic
fluorophores and; .gamma.=1 for linear fluorophores. Calculations
were performed using the NIRFAST package implementing the finite
element method. W was factorized according to (4); W=U.SIGMA.V*,
(4) where U and V are unitary matrices containing the left and
right singular vectors of W, and; .SIGMA. is a diagonal matrix
containing the singular values of W. The column-space of V is
spanned by the image-space modes, while the column-space of U is
spanned by the detection-space modes. The singular values of W
denote how effectively a given image-space mode can be detected by
an experimental setup [Ref. 2].
FIG. 6 shows the normalized singular-value distribution of W. The
x-axis shows the singular value index (1120) and the y-axis shows
the normalized singular value intensity (1121). For clarity, only
every second singular value are shown. The cross (1122) and plus
(1124) signs represent the linear fluorophore (.gamma.=1), the
former for the single-beam excitation (1122), while the latter for
the combined single-beam excitation and dual-beam excitation
(1124). As seen, the normalized intensities of the additional
singular values due to dual-beam excitation (1124) have dropped to
machine precision, which indicates that the measurements with
dual-beam excitation may not alleviate the ill-posedness of FDOT.
In other words, the type-D images may not provide more information
than the existing type-S images. Hence, it may not improve the
quality of the reconstructions. However, for the quadratic
fluorophore (denoted by asterisk (1123) and dot (1125) signs in
FIG. 6, the intensities of the additional singular values (1125)
are still significant. This implies that type-D images will
contribute to the quality of the reconstructions.
A single excitation beam may first be used to scan over a
(3.times.3) grid, and capturing one image for each scanned position
by a CCD camera. In the next step, two excitation beams, located at
two nearest-neighboring sites of the same grid, can be
simultaneously employed to illuminate the phantom, giving 6 extra
type-D images.
FIGS. 7A-7B shows the three-dimensional rendering of the
reconstructed upconverting nanoparticles. The red cylinders in the
subfigures are identical and represent the true fluorescent
lesions. In the reconstruction of FIG. 7A, only type-S images were
used. As can be seen, the shape of the fluorescent lesion is
overestimated. This overestimation may be explained by the
ill-posedness of the inverse problem. When adding type-D images,
the reconstruction of the fluorescent lesion shape is improved
remarkably, as shown in FIG. 7B. Images of type D contribute to the
inverse problem and lead to better reconstructions for the
quadratic upconverting nanoparticles.
It is disclosed an additional unique advantage of the nonlinear
power dependence of upconverting nanoparticles. This advantage
enables the possibility to obtain additional information for the
inverse problem by using images taken with two or more excitation
beams simultaneously. The same advantage could not be found when
using linear fluorophores, e.g., Rhodamine 6G.
Enhancing Upconversion Emission by Pulsed Excitation
As shown in FIG. 8, illustrating upconversion spectrum for
NaYF4:Yb3+, Tm3+, upconverting nanoparticles can emit emission
bands in the near infrared (.about.800 nm), red (.about.648 nm) and
blue (.about.475 nm) ranges under excitation of 975 nm.
As described above, the intensities of these upconverting emission
bands have nonlinear dependencies on excitation intensity. The
dependence in low intensity range can be described by
I.sub.f=kI.sub.ex.sup.n (5)
where I.sub.f is the upconversion fluorescence intensity; k is a
constant; I.sub.ex is the excitation intensity; n is the number of
excitation photons required in order to generate one emission
photon.
The power dependencies of the near infrared, red and blue emission
bands are shown in FIG. 9, illustrating power dependencies of near
infrared, red and blue upconversion emission bands of NaYF4:Yb3+,
Tm3+ nanoparticles under excitation of 975 nm.
Quantum yield is defined as the ratio between the numbers of
emitted photons and the number of absorbed excitation photons.
Because of their nonlinear power dependencies shown in FIG. 9,
upconversion emissions have power-density dependent quantum yields
instead of constant quantum yields, as illustrated in FIG. 10,
showing quantum yields of near infrared, blue, red upconversion
emission bands at various power densities.
A method 100 of imaging a region in a scattering medium by diffuse
luminescence molecular imaging according to an embodiment of the
invention comprises (FIG. 28) providing 101 at least one non-linear
luminescent marker in a scattering medium at a marker position in
said region, exciting 103 the non-linear luminescent marker by
excitation light emitted by one or more light sources into an
excitation volume from at least one light source position, and
detecting 107 luminescence from the luminescent marker due to the
excitation light by a detector at a luminescent light detection
position, wherein the excitation light comprises pulsed excitation
light.
The quantum yield increases with power density and gradually
approach a constant. A gain in the signal level is provided from
pulse excitation compared to continuous wave excitation with the
same average power, because the pulse excitation has higher peak
power.
Confining the same number of excitation photons in a narrow time
window through pulse excitation can hence be provided for more
efficiently using excitation photons in order to get stronger
upconversion emission light. This is confirmed in FIG. 11, showing
spectra of NaYF4:Yb3+,Tm3+ nanoparticles using CW or pulsed light
with the same average power as excitation. The pulse has a pulse
width of 10 ms and a period of 100 ms, and the beam sizes both for
the CW and pulse excitation is 0.70 mm in diameter in FIG. 11. As
shown in FIG. 11, a signal gain by a factor of around 2 is obtained
by using the pulse excitation compared with the CW excitation with
the same average power of 1.94 mW.
By having pulsed excitation light a significant increase in quantum
yield when using upconverting nanoparticles is accordingly
provided. Further, pulsed excitation light provides for single-shot
deep tissue imaging, large imaging depths and short data
acquisition times compared with continuous wave excitation. Thermal
side effects of the excitation light are also suppressed because of
the pulsed light.
Pulsed excitation light also provides for diffuse optical imaging,
photodynamic therapy and remote activation of biomolecules in deep
tissues. The aforementioned effects have been described in more
detail under "Single shot imaging" below, which is part of the
present application.
An additional advantage with pulsed excitation and UCNPs with long
emission lifetime is that it is possible to suppress scattered
excitation light by employing delayed detection. This has
previously not been utilized for UCNPs. For macroscopic imaging
inside tissue a great advantage of the UCNPs is the anti-Stokes
shift of the light emission, proving means to suppress the tissue
autofluorescence. This is known and provides in theory a total
background free signal, of great interest. Even though tissue
autofluorescence can be totally suppressed, there is still in
practice an issue in prior art with spectrally filtering out the
signal from the much stronger scattered excitation light. With
pulsed excitation and time-delayed detection, this suppression
would be more efficient, and the advantage with total
background-free signal would be easier to utilize in practice with
pulsed excitation. The method 100 may thus comprise the step of
time-delaying 105 the detection of the luminescence to provide for
detection of a signal without the influence of the excitation light
scattered in the medium. FIG. 21 shows an example where an
excitation pulse 201 with length (w) is followed by luminescence
with a decaying intensity during a time interval 210. The method
may thus comprise the step of detecting 108 the luminescence during
the time interval 210 succeeding the pulse 201 of said excitation
light. The time interval may be in the range of 1-100 ms. The
system 600 may accordingly comprise a detector unit 601 that is
operable to detect the luminescence during a time interval 210
succeeding a pulse of said excitation light.
Pulse Width Dependent Gain:
Under CW excitation, if the excitation power is doubled, the
fluorescence intensity will be four times higher if the
power-densities are in the non-saturation power-density regime, due
to the quadratic power dependence (case 1).
Under pulsed excitation, it will take a certain time (determined by
the lifetimes of intermediate energy levels) to reach steady state.
During the rise time, the fluorescence intensity is weaker than
that at steady state condition. Thus, comparing the fluorescence
intensities under CW excitation and under a square-wave pulse
excitation with twofold higher peak power, during the period of
pulse duration, the latter will not be fourfold higher but less
than the former, which is different from case 1. Hence, the gain in
upconversion emission intensity by pulse dexcitation is pulse width
dependent. If the pulse width is too short, the upconversion system
will be far away from steady state during the pulse duration, thus
the gain will be smaller or no gain at all. The pulse width should
be long enough, and it can be determined with the assistance of the
observation of the upconversion spectra under CW excitation and
pulse excitation with different pulse width. If the pulse width is
long enough to reach steady state, the normalized spectrum under
pulse excitation should adequately approach that under CW
excitation. FIG. 12 shows the upconversion emission spectra
(normalized at 800 nm) under CW excitation and pulse excitation
with different pulse width (1 ms, 5 ms, 10 ms). The CW excitation
has an average power of 24 mW (beam size 0.70 mm). The peak powers
of all the pulses is 24 mW (beam size 0.70 mm). The inset of FIG.
12 shows the zoomed-in part of the range of 434-674 nm. As seen,
when the pulse width is 10 ms, the difference between the spectra
from CW and pulse excitation is less than 10%, indicating steady
state or quasi steady state is reached.
When the pulse width is increased from 5 ms to 10 ms, the gain is
increased, as shown in FIG. 13 (showing upconversion spectra under
excitation of CW and pulses with the same average power of 31.04
mW) and FIG. 14 (showing the gain of signal at 800 nm under pulse
excitation with 5 and 10 ms pulse widths under different average
excitation power).
Power Dependent Gain:
The quantum yield of the emission at 800 nm gradually approaches a
constant when increasing the power density, as shown in FIG. 10.
Thus, the gain decreases with increasing the average power. When
the average power is in the range in which quantum yield is a
constant, no gain any more. We even lose some signal due to the
quantum yield loss during the rise time above mentioned, as shown
in FIGS. 15a-f, 16.
Since the blue and red emission have large slopes in the power
dependence curves (as shown in FIG. 9), e.g., larger n in equation
(5), thus higher order power-density dependent quantum yields than
800 nm emission (as shown in FIG. 10), and they are more difficult
to get saturated (as shown in FIG. 9 and FIG. 10), in stark
contrast their gain are larger than that of 800 nm. Even at the
maximum power investigated, there are still gains by factors of 5.3
and 5.7 for the red and blue emissions, respectively, as shown in
FIGS. 17a-c, 18a-b.
It is noteworthy to point out that the gain by pulse excitation is
related with the parameters of the pulse. All the above results are
obtained with square wave pulse excitation. Different pulses such
as triangle or sine wave will give different results, but signal
gain can be also expected. The duty cycle of the square wave is
another key parameter, which determines what maximum gain could be
obtained. The pulse with a pulse width of 10 ms and period of 100
ms has a duty cycle of 10%, so the maximum gain could be 10 (1/duty
cycle) for 800 nm emission. Examples show a gain by a factor of
around 3.8. By using smaller duty cycles, larger gain may be shown.
Hence, the present disclosure provides for improved gain by using
small duty cycles, for example well below 50% duty cycle which
would only allow a gain by a factor of 2. By using small duty
cycles it is provided for achieving optimally high peak power for
improved imaging abilities with the advantages described herein. By
using single pulse excitation as explained below even higher power
density can be achieved in order to exploit high intrinsic QY of
upconversion nanoparticles. The pulsed excitation provides
accordingly for delivering high power densities while complying
with ANSI standards.
Power-Scanning Tomography
The change of the power dependence shown in FIG. 9 can be used to
perform power-scanning tomography using a single excitation point
and any number of detection points. The concept can be briefly
summarized as a discretization of the power-dependence curve, where
at each discretized region, a given slope coefficient is used as
input to generate (simulate) the expected fluorescence. This can be
further used to perform a tomographic reconstruction using
conventional optimization methods in an extremely fast fashion, by
only power-scanning the excitation source with no spatial scanning.
Advantages include, speed, no-moving parts, and simplified
instrumentations.
In conclusion, by using pulse excitation, upconversion emission
intensity can be enhanced compared with CW excitation with the same
average power. The enhancement originates from the use of the same
amount of excitation photons with a higher efficiency, which
results from the power-density dependent quantum yield of
upconversion nanoparticles, here NaYF4:Yb3+,Tm3+. The gain is pulse
width and power dependent.
This proposed technique is a general approach for utilizing the
upconversion capability more efficiently. It will work for not only
Yb3+/Tm3+ codoped upconverting nanoparticles, but also for any
upconverting nano- or bulk-materials. It works even better for high
order upconversion emission, such as the blue and red emission of
Tm3+ from three-photon processes. This approach can be useful in
enhancing shorter wavelength upconversion emission needed for
photodynamic therapy in biological tissue.
The power-dependence feature of upconversion emission can be used
to perform power-scanning tomography using a single excitation
point.
Upconversion Signal Enhancement by Pulse Excitation in Tissue
Phantom
The validity of this technique is also confirmed by measurement in
tissue phantom, see FIGS. 19a-c. The experiments were carried out
in a 20 mm thick liquid tissue phantom with a reduced scattering
coefficient of 10 cm.sup.-1 and an absorption coefficient of 0.5
cm.sup.-1, made of water, intralipid and ink. A glass tube with
inner diameter of 2 mm, filled with hexane colloidal of
NaYF4:Yb3+,Tm3+ nanoparticles (c=1 wt %), was inserted into the
phantom with a depth of 10 mm. Two laser sources were used for the
comparison, a CW laser diode at 975 nm and a pulse laser with
tunable pulse width and period at the same wavelength. The spot
sizes of the lasers were 1 mm in diameter. Two different settings
were used for the pulse laser: (a) Setting 1: 5 ms pulse width, 250
ms period, FIG. 19b; (b) Setting 2: 10 ms pulse width, 500 ms
period, FIG. 28c. The upconversion emission images taken at 800 nm
under the excitation of CW or pulse excitation are shown in FIGS.
19a-c. The average power was kept the same (100 mW) for all the
measurements. The intensities under pulse excitation is around 6
and 6.75 times higher than that under CW excitation for Setting 1
and 2, respectively.
FIG. 19a show an image taken at 800 nm under the excitation of a CW
laser diode with a power of 100 mW.
FIG. 19b show an image taken at 800 nm under the excitation of a
pulse laser (square wave, pulse width 5 ms, period 250 ms) with an
average power of 100 mW.
FIG. 19c show an image taken at 800 nm under the excitation of a
pulse laser (square wave, pulse width 10 ms, period 500 ms) with an
average power of 100 mW.
The higher intensities obtained from the pulsed excitation provides
for improved imaging due to the increase of the upconversion signal
level. Further, the pulsed excitation reduce the heating effects in
the biological tissue, while maintaining the increased signal level
and improved imaging. For example, a single shot (10 ms-100 ms) by
a laser (peak power up to e.g., 100 W) to generate a strong peak
signal, and then turning off the excitation source will allow the
biological tissue to cool down, in order not to overheat the tissue
but dramatically increase the emission signal. Further, it would be
possible to use a very low-power light source with pulsed
excitation light to achieve acceptable signal levels for the
imaging, in comparison to continuous wave laser diode that would
require more power to produce the same result.
Single-Shot Imaging with Pulsed Excitation Light
The limited quantum yield (QY) of upconverting nanoparticles
(UCNPs), especially at low light conditions, is of major concern
for most potential biological applications. Two highly potent
techniques in the field are deep tissue optical imaging and
photodynamic therapy (PDT), which both require high QY. The present
low QY issue hinders the potential of these techniques by resulting
into increased treatment and data acquisition times and shallow
applicable depths. Although, the low QY can to some extent be
overcome by elevating the excitation light level, such improvements
are restricted for CW excitation by risks of side-effects in terms
of tissue heating (regulated by the ANSI standards). According to
embodiments of the invention, by employing pulsed excitation, it is
provided for to break through the low power-density limit of
upconversion (UC) emission while limiting the thermal effect of the
excitation light. In addition, the applicability of UCNPs may be
further boosted by utilizing single-shot excitation schemes.
Similar to multiphoton microscopy, pulsed excitation may provide
high photon density during the pulse, while keeping the average
power (meaning the deposited energy responsible for the heating)
moderate. Due to the nonlinear power-density dependence of UC
emission, pulsed excitation provides for beneficial effects as
discussed in this disclosure.
Examples of the present disclosure take excitation dynamics of UC
emission into account to overcome issues with previous techniques
that demonstrate low quantum yield. The below disclosure gives
examples of experiments and simulations demonstrating significant
QY increase which can be achieved by using pulsed excitation light
in a method, system and use of a system according to embodiments of
the invention. E.g. pulsed excitation light is used with matched
pulse characteristics, i.e., with sufficiently long pulse width and
non-saturated transitions to provide for the advantageous effects.
This makes pulsed excitation an ideal excitation approach for
UCNPs, especially for deeply located tissue volumes. In addition,
single-shot imaging of UCNPs can be implemented due to the
increased QY, in which the data acquisition time can be shortened
by orders of magnitude while improving the imaging depth as
compared to CW light excitation causing the same temperature
increase. Thus the present disclosure has the potential to
fundamentally broaden the applicability of UCNPs in deep tissue
regions relying on diffuse light excitation.
The excitation dynamics can be modelled using rate equations.
Without loss of generality, NIR UC emission at 800 nm of Yb3+/Tm3+
codoped system may be used as a model in the below example. FIG.
29(a) shows the simulated QY at steady state conditions following
CW excitation of different power densities. As seen, the QY
increases with the excitation power-density in a complex rather
than a purely linear manner, and exhibits a feature of gradual
saturation, i.e., approaching a constant at high excitation
power-densities. FIG. 29(b) presents the simulated temporally
cumulative QY under CW excitation and under pulsed excitation in
the first pulse period. The CW excitation has a constant
power-density of 1 W/cm2. The pulsed excitation, having a frequency
of 2 Hz and a duty cycle of 4%, has power-densities of 25 W/cm2 and
0 W/cm2 at the "on" and "off" states, respectively, thus resulting
in the same average power-density as the CW excitation. As seen,
under CW excitation, the UC emission has a constant QY except at
the very early stage when the energy levels are populated due to
transient effects of the excitation. This constant QY is associated
with the steady state of the UC system, and given by the QY at the
power-density of 1 W/cm2 in FIG. 2(a). Under the pulsed excitation,
the QY is very small at the start of the laser pulse, and then
increases with time. If the length of pulse duration allows, the QY
will surpass the QY under the CW excitation, and asymptotically
approach a maximum. This maximum is restricted to the QY at steady
state at the power-density of 25 W/cm2 in FIG. 29(a). Clearly, the
advantage of using pulsed excitation to replace the equivalent CW
excitation is that the late excitation photons can be potentially
used with higher energy conversion efficiency, while the cost is
that the early excitation photons in each pulse period are used
with lower efficiency than in the CW excitation. Through balancing
the increased power-density and decreased excitation time under the
same amount of energy, an overall UC signal gain can be
expected.
Long-term QY in multiple periods under pulsed excitation was
investigated, in order to determine the influence of the pulse
width on the potential signal gain. The average power-density was
kept at 0.1 W/cm2. The pulsed excitation used in this study had the
same duty cycle of 4% unless otherwise specified, and its frequency
was adjusted in order to achieve different pulse widths. As
illustrated in FIG. 29(c), a significant UC signal gain is obtained
by using pulsed excitation when the frequency is well below 50 Hz.
For example, the signal gain by the 2-Hz square wave in the time
interval of [0, 500] ms is approximately 8. The signal gain
decreases with frequency, i.e., increases with pulse width. When
the frequency is even higher, e.g., up to 100 Hz, the signal
generated by the pulsed excitation becomes slightly smaller than
that generated by equivalent CW excitation. It should be noted that
the signal gain decreases with the applied power-density. When the
average power-density is increased to 1 W/cm2, the gain declines to
2. This can be ascribed to the gradual saturation property of UC
emission, as indicated in FIG. 29(a).
In order to experimentally validate the gain in signal due to
pulsed excitation as indicated by simulations above, experiments
were carried out on colloidal stable stable core-shell
NaYF4:Yb3+,Tm3+@NaYF4 UCNPs dispersed in hexane. The prepared UCNPs
emit the major UC emission bands at around 800 nm under excitation
of 975-nm light, as shown in FIG. 3(b). The intensities of this NIR
UC emission under CW excitation and under pulsed excitations with
different pulse widths were recorded. The average power-density of
the excitation light was kept at 0.12 W/cm2. As shown in FIG.
26(a), a signal gain, monotonically increasing with pulse width,
was obtained by using the pulsed excitation even with a pulse
duration of 0.8 ms. When the pulse width reaches 20 ms, the gain is
as high as 8.7. These results agree well with the simulated results
presented in FIG. 29. It is noteworthy to point out that the
required pulse width for signal gain in the present case
(.about.0.8 ms) is much shorter than the rise time of the UC
emission, approximately 10 ms as shown in the inset of FIG.
26(a).
The dependence of the UC signal gain on the applied power-density
was also investigated using a square-wave excitation with a fixed
pulse width of 20 ms and a period of 500 ms, together with the
equivalent CW excitation. FIG. 26(b) shows the UC signal gain by
the pulsed excitation at various average excitation
power-densities, where a decreasing trend with increasing
excitation power densities is clearly seen. At the minimum
power-density investigated (.about.0.12 W/cm2), the signal gain is
approximately 8.6, while at the maximum power-density (.about.4.65
W/cm2), the UC signal generated by the pulsed excitation is
slightly weaker than that generated by the CW excitation. The UC
emission intensity dependence on the excitation power-density
exhibits a smaller slope than under the CW excitation, as shown in
the inset of FIG. 26(b). This can explain the signal-gain trend
above. The amplification effect of increasing the excitation
power-density here essentially originates from the non-linear
power-density dependence of the UC emission. Thus, a higher-order
power-density dependence would result in a larger UC signal gain.
This is confirmed by the measurements on the blue and red UC
emissions, both generated through a three-photon excitation
process. They exhibit significantly larger signal gains than the
NIR UC emission at any given average power-density, as shown in
FIG. 30.
The merit of using pulsed source as the excitation approach to
image deeply located UCNPs was subsequently validated in a liquid
tissue phantom. The phantom, made of water, intralipid and ink, was
determined by a photon time-of-flight spectroscopy (pTOFS) system
to have a reduced scattering coefficient of .mu.'.sub.s=10.1
cm.sup.-1 and an absorption coefficient of .mu..sub.a=0.52
cm.sup.-1 at 975 nm, and had a thickness of 17 mm. A glass tube
with an inner diameter of 2 mm, containing the colloidal core shell
UCNPs (c=1 wt %), was inserted into the phantom as the luminescent
inclusion to mimic a UCNP-labeled target (e.g., a tumor) inside
real tissue. One out of two 975-nm lasers, including a CW laser
diode and a pulsed laser with a pulse width of 20 ms and a period
of 500 ms, was used to provide the excitation light. The average
power-density impinging on the surface of the tissue phantom was
1.2 W/cm2 for both excitation approaches. The excitation source and
the detector were positioned in a trans-illumination geometry.
When buried at a depth of 10 mm from the source, the luminescent
inclusion was barely detectable under CW excitation even with an
exposure time of 10 s, as shown in FIG. 27(a), whereas by using
pulsed excitation, the signal-to-background ratio was significantly
increased by a factor of approximately 7 under the same detection
conditions, as illustrated in FIG. 27(b). An obvious implication is
that the data acquisition time can be remarkably reduced and the
imaging depth can be increased if keeping the signal quality as the
equivalent CW excitation. The QY of UC emission may be further
optimized by using a single pulse as excitation providing even
higher power-density. For example, the maximum permissible
power-density for exposure to human skin at 975 nm is 17.4
W/cm.sup.2 for a repetitive pulse excitation with a pulse width of
20 ms and a frequency of 2 Hz, while the number for a 50-ms single
pulse is as high as 36.9 W/cm.sup.2, referring to the supplemental
material. Such strong single pulse with a pulse width longer than
the rise time of the UC emission enables the UCNPs to be used in a
very efficient way in terms of energy conversion. This excitation
approach would improve the imaging ability of using UCNPs without
violating the ANSI standard.
The feasibility of single-shot imaging was thus experimentally
investigated. A 50-ms single pulse providing an excitation
power-density of 36.9 W/cm.sup.2 was used. When the luminescent
inclusion was placed at a depth of 13 mm from the source, it could
still be relatively well detected using the single pulse excitation
with an exposure time of 1 s, even using an epi-fluorescence
imaging setup, as shown in FIG. 27(d). Nevertheless, when the CW
laser was used as the excitation source, also outputing the maximum
permissible powerdensity by ANSI standard on the same illumination
area, i.e., 709.6 mW/cm2, the inclusion was not detectable at all
even with a much longer exposure time of 10 s, as shown in FIG.
27(c). The exposure time for the single pulse excitation may be
shortened to 50 ms still without loss in the UC signal quality, as
long as the excitation source and the detector are synchronized.
The examples demonstrated here show the great potential of
single-shot UCNP excitation in UCNP-guided deep tissue optical
imaging.
Single-shot imaging of UCNPs in deep tissue phantom can thus be
accomplished according to embodiments of the invention, by
employing pulsed excitation to significant increase the QY. The
pulsed excitation approach thereby greatly increase the
applicability of UCNPs not only in diffuse optical imaging but also
in many other biomedical applications, such as photodynamic therapy
and remote activation of biomolecules in deep tissues. Further,
metallic nanostructures may be effective in enhancing UC emissions
owing to their local field enhancement effect by surface plasmonic
coupling. Combining pulsed excitation and the decoration with
metallic nanostructure may therefore allow a major scheme of using
UCNPs in the diffuse light regime, due to the synergistic effect in
increasing the excitation power-density. A method 100 according to
an embodiment of the invention may thus comprise the step of
providing 117 metallic nanostructures at said medium for exposure
to said pulsed excitation light The pulsed excitation approach will
also increase the applicability of migration-mediated UC emissions
from ions such as Eu3+ and Tb3+ in biological applications, due to
their high-order multi-stepwise excitation nature via excited Tm3+.
In addition, this disclosure provides a general method for
promoting the applications of nonlinear fluorophores (including
UCNPs and triplet-triplet annihilation based upconverters) at low
light conditions by increasing the excitation fluence rate through
a limited illumination area.
The method 100 may thus comprise the step of exciting 104 the
non-linear luminescent marker with a first pulse 201, i.e. the
pulsed excitation light comprises at least one pulse of light, and
further the step of detecting 106 luminescence from the luminescent
marker due to said excitation light from said first pulse for
providing single pulse luminescence molecular imaging from the
first pulse. This single pulse imaging provides for several of the
above described advantages over CW excitation.
The method 100 may comprise the step of matching 102 pulse
characteristics of the at least one pulse, such as the length (w)
of the pulse, with energy level transitions conditions of the
non-linear luminescent marker to substantially provide for a
desired population of energy levels of said non-linear luminescent
marker related to emission of upconverted light so that said
upconverted light is produced in a very efficient manner. The
dynamics of the energy level transitions involved in the
excitation/emission process is thereby taken into account to adapt
the characteristics of the pulse. I.e. in order to provide for
adequate and optimized intensity of the luminescence the pulse
characteristics can be tailored to provide for the particular
conditions by which population of the energy levels follows the
desired scheme, e.g. by taking into account the duration of the
lifetimes of the excited states that are involved in the emission
process. The method 100 may thus comprise determining 116 a pulse
width and/or a pulse waveform of said pulsed excitation light to
provide excitation of said non-linear luminescent marker.
In this context, the method 100 may comprise the step of
determining 104 a pulse length (w) of the pulsed excitation light
to be in a range that provides excitation to the energy levels
involved in the emission of upconverted light. The length of the
pulse may be determined based on calculation of the lifetimes of
the energy levels. The length of each pulse in a train of pulses
may be in the range of about 1-100 ms. The system 600 may thus
comprise a processing unit 603 operable to determine a pulse length
of said excitation light based on calculation of energy level
transitions conditions of said non-linear luminescent marker such
as life time calculations.
The method 100 may comprise the step of determining 104 a pulse
length (w) of the pulsed excitation light to be in the range of
about 20-200 ms for single pulse luminescence molecular imaging,
such as described in relation to FIG. 27d.
The method 100 may comprising determining 115 a dependence of the
detected luminescence on the power of said excitation light for
setting a predetermined characteristic of said pulsed excitation
light.
Referring to FIG. 26a, signal gain may be obtained by using a pulse
width as short as 0.8 ms. In addition, the optimal pulse width is
material dependent. For instance, the optimal duration for
Yb3+/Er3+ codoped materials would be shorter than that for
Yb3+/Tm3+ codoped materials, as the intermediate energy states of
Er3+ ions have usually short lifetimes than those of Tm3+ ions.
Having a pulse width of about 100 us will typically not provide
sufficient signal gain.
A pulse length longer than the associated lifetimes may be
advantageous in providing improved quantum yield and thereby
improving the imaging capabilities. Gain may still be provided by
having a pulse length of 0.8 ms while having lifetimes that are
more than 0.8 ms. The gain by using pulsed excitation essentially
originates from the higher peak power density. Too short pulse
length will "eat" the benefit brought by the higher peak power
density. When using 100 Hz, 50% duty cycle pulse, corresponding to
0.4 ms pulse width, there is no signal gain by using pulsed
excitation for specific nanoparticles. Femtosecond or microsecond
pulse laser to excite upconverting nanoparticles does not show
signal gain as provided by the present disclosure. For the upper
limit of the pulse width, the concern is that the detection system
would wait too long time if the pulse length is large (since
generally the duty cycle is small for this technique). For
instance, for a pulsed excitation with a 100 ms pulse width and 10%
duty cycle, the waiting time before collecting next-period
luminescence signal generated by the laser is 900 ms. It is
acceptable. But if it is even longer, it would not be economic in
time in experiment.
For single pulse excitation it may be advantageous to increase the
pulse length. The concern is that the laser will have a rest for
quite some time after delivering the single pulse, so the
luminescence signal will be generated only in such an interval. If
the pulse duration is too short, the generated emission photons
escaping from the surface of tissue would be too few to give a good
signal (but QY increase is still there compared with equivalent CW
excitation). A preferable interval of duration for single pulse
could be 20-200 ms. If the single pulse is too long, it will too
much like a CW source. In that case, it is not allowed to use
preferably high peak power density according to ANSI standard.
Reference is now made to FIGS. 22-25. The method 100 may comprise
the step of varying 109 the power density (I) of the pulsed
excitation light as a function of time (t) (as illustrated in FIG.
23 with reference to power density curves 206, 207), determining
110 a quantum yield dependence (Q/I) of the luminescence on said
power density, and determining 111 a relative depth coordinate 203
of the marker position in the scattering medium based on the
quantum yield dependence. A series of pulses (w1, w2), such as
illustrated in FIG. 20 with different power densities may be used
as excitation light. The quantum yield can be calculated for each
of the pulses, and a dependence (Q/I) such as illustrated in FIG.
25 can be determined. Depending on how the quantum yield varies as
the power density of the excitation light is varied, it is possible
to determine if the marker is located relatively deep or shallow in
the medium, as illustrated for markers 209 and 208 in FIG. 22
respectively. A marker 208 which has a more shallow position 204 in
the medium will exhibit less variance in the luminescence quantum
yield as the power density is varied, compared to a marker 209
which has a deeper position 205, as indicated in the Q/I curve in
FIG. 25. The method 100 may comprise determining 112 the relative
depth coordinate 203 based on a derivative (dQ/dI) of the quantum
yield dependence. And the method 100 may accordingly comprise
exciting 113 in sequence the non-linear luminescent marker with a
first and second pulse 201, 202, having first and second power
densities (I.sub.1, I.sub.2) respectively, and determining 114 the
relative depth coordinate 203 based on a variation in the quantum
yield from the first and second pulses. This allows for
distinguishing between makers at different depths, without having
to take into account the influence of their relative sizes. The
system 600 may thus comprise a control unit 602 that is operable to
vary the power density of the pulsed excitation light as a function
of time (t), and a second processing unit 604 that is operable to
determine a quantum yield dependence (Q/I) of the luminescence on
the power density, and to determine a relative depth coordinate 203
of the marker position in the scattering medium based on said
quantum yield dependence.
The system 600 may comprise a control unit 605 for performing the
method 100 as described above. Further, use of a system 600 for
performing the method 100 is provided according to the present
disclosure. More particularly, the use of a system 600 is disclosed
for luminescence imaging or luminescent tomography of tablets,
and/or for diffuse optical imaging, and/or photodynamic therapy
and/or remote activation of biomolecules in deep tissues, and/or
single-shot deep tissue imaging, and/or for in-vivo or in-vitro
luminescence imaging or luminescent tomography of a small animal,
and/or for functional diagnostics, such as cancer diagnostics, by
said luminescence imaging or luminescent tomography, and/or for
superresolution microscopy comprising stimulated emission depletion
(STED) or single-molecule detection using said non-linear
luminescent marker as probe.
Potential Use of Pulsed Excitation in Upconverting Nanoparticles
Based Photodynamic Therapy
This proposed technique is a general approach for utilizing the
upconversion capability more efficiently. It will work for not only
Yb3+/Tm3+ codoped upconverting nanoparticles, but also for any
upconverting nano- or bulk-materials. It works even better for high
order upconversion emission, such as the blue and red emission of
Tm3+ from three-photon processes. This approach is in particular
useful in enhancing shorter wavelength upconversion emissions which
are needed for upconverting nanoparticles based photodynamic
therapy in biological tissue.
The present invention has been described above with reference to
specific embodiments. However, other embodiments than the above
described are equally possible within the scope of the invention.
The different features and steps of the invention may be combined
in other combinations than those described. The scope of the
invention is only limited by the appended patent claims.
The method may be performed in-vivo at a living human or animal
body. In this case, the markers may be preintroduced into the body
in any manner, such as by injection into the blood stream or
subcutaneously or directly into a tumour, or alternatively by
topical application, pulmonary and other non-invasive methods. Such
preintroduction can be performed separately from the remaining
method. Such preintroduction can be performed in connection with
the remaining method but shortly before.
Alternatively or additionally, the method may be performed at a
human or animal body, which is sacrificed after the method is
performed.
Alternatively or additionally, the method may be performed in vitro
at a non-living human or animal body or part of a body, for example
a brain-dead human or animal body.
Alternatively or additionally, the method may be performed at
non-medical fields, such as filters or tablets.
Superresolution Microscopy Using UCNP as Probes
Superresolution microscopy has recently been developed and become a
very interesting and useful tool for much biological research.
There are two types of superresolution microscopy, one that relies
on non-linear optical effects and one on single molecule detection.
They both have in common that selected molecules provide a signal,
while others are filtered out. The first category include
stimulated emission depletion (STED) and saturated structured
illumination microscopy (SSIM), while the single molecule detection
comprises PALM (photoactivated localization microscopy), FPALM
(fluorescence photoactivated localization microscopy) and STORM
(stochastic optical reconstruction microscopy). The first category
of superresolution microscopy utilize that probe can emit light,
while nearby probes can be made non-emitting. The excited state
will be depopulated for theses nearby probes. This procedure sets
requirements for the probe used, one is that it has to be extremely
photostable (as this is a non-linear effect requiring relatively
high excitation power), non-blinking (as they should be active all
the time while in the active state) and should have several energy
levels that the probe can be light-switched to. UCNP could be an
ideal probe for STED, with the unique properties, fulfilling these
requirements.
The other category relies on that the probes can be photoswitched
to other energy levels and become inactive. This is a
single-molecule regime with low light levels. It could also rely on
spontaneous photo-blinking. UCNP could become interesting probes
for these techniques with the many energy levels in these
probes.
The following references are incorporated by reference herein in
their entirety for all purposes: [Ref. 1] C. T. Xu, J. Axelsson,
and S. Andersson-Engels, Appl. Phys. Lett. 94, 251107 (2009). [Ref.
2] J. P. Culver, V. Ntziachristos, M. J. Holboke, and A. G. Yodh,
Opt. Lett. 26, 701 (2001). [Ref. 3] E. Alerstam, S.
Andersson-Engels, and T. Svensson, J. Biomed. Opt. 13, 041304
(2008).
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